WO2005011502A1 - Tomographe - Google Patents

Tomographe Download PDF

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Publication number
WO2005011502A1
WO2005011502A1 PCT/JP2004/008379 JP2004008379W WO2005011502A1 WO 2005011502 A1 WO2005011502 A1 WO 2005011502A1 JP 2004008379 W JP2004008379 W JP 2004008379W WO 2005011502 A1 WO2005011502 A1 WO 2005011502A1
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WIPO (PCT)
Prior art keywords
image
phantom
radiation
subject
imaging system
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PCT/JP2004/008379
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English (en)
Japanese (ja)
Inventor
Hironori Ueki
Yasutaka Konno
Original Assignee
Hitachi Medical Corporation
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Publication date
Application filed by Hitachi Medical Corporation filed Critical Hitachi Medical Corporation
Priority to US10/566,205 priority Critical patent/US20070116183A1/en
Priority to JP2005512455A priority patent/JPWO2005011502A1/ja
Publication of WO2005011502A1 publication Critical patent/WO2005011502A1/fr

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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/58—Testing, adjusting or calibrating thereof
    • A61B6/582—Calibration
    • A61B6/583—Calibration using calibration phantoms
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/52—Devices using data or image processing specially adapted for radiation diagnosis
    • A61B6/5258—Devices using data or image processing specially adapted for radiation diagnosis involving detection or reduction of artifacts or noise

Definitions

  • the present invention relates to a radiation tomography apparatus, and in particular, reduces an artifact in a radiation tomographic image generated due to a change in radiation quality of a radiation in a subject or a nonlinearity of input / output characteristics of a radiation detector.
  • a radiation tomography apparatus reduces an artifact in a radiation tomographic image generated due to a change in radiation quality of a radiation in a subject or a nonlinearity of input / output characteristics of a radiation detector.
  • artifacts may occur in a reconstructed tomographic image of a subject (hereinafter referred to as a reconstructed image).
  • the main types of such artifacts are ring artifacts, which appear as a ring pattern in the reconstructed image, and dark band artifacts, which appear as a black band pattern between high radiation absorbing substances. is there.
  • the main causes of ring artifacts are that the sensitivity of the radiation detector varies from pixel to pixel, and that the input / output characteristics of the radiation detector are not ideal and have nonlinearities.
  • the main cause of dark band artifacts is the fact that the radiation quality changes in the subject (beam hard Jung effect).
  • the most common artifact reduction method is the air calibration method.
  • the air calibration method an air image taken without placing a subject in advance is prepared.
  • the signal intensity distribution of the air image corresponds to the product of the intensity distribution of the radiation incident on the detector and the detector sensitivity distribution. Therefore, in each pixel of the detector, the signal of the photographed image is converted into the signal of the air image. By the division, the dispersion of the radiation intensity distribution and the detector sensitivity can be corrected.
  • the water correction method is an extension of the air calibration method, and uses a water image obtained by photographing a water bottle phantom having a cylindrical or elliptical cylindrical shape instead of an air image.
  • a phantom calibration method has been proposed (see, for example, Japanese Patent Publication No. Sho 61-54412).
  • the phantom calibration method is a method of correcting photographing data of a subject based on a conversion function created in advance.
  • the conversion function is a polynomial for converting the measured value of the captured image into its theoretical value, and the relationship between the measured value and the theoretical value is derived in advance using a calibration phantom.
  • the air calibration method can correct the intensity distribution of the radiation incident on the radiation detector and the sensitivity variation of the radiation detector, the ringer chip act is greatly reduced.
  • the nonlinear input / output characteristics of the radiation detector cannot be corrected, there is a problem that the ring artifact cannot be completely removed. Also, due to beam hard Jung There is a problem that the band artefact cannot be removed.
  • the water correction method has an advantage that it can remove ring artifacts and dark band artifacts due to the nonlinear input / output characteristics of the detector, in addition to the capturing effect of the air calibration method.
  • the radiation absorption amount of the water bottle phantom and the object cannot be completely matched, there is a problem that the above-described removal accuracy is low.
  • the phantom calibration method can calibrate the signal strength of the captured image over a wide range of the detector dynamic range. For this reason, there is a merit that the ring artifact and dark pan artifact caused by the nonlinear input / output characteristics of the detector can be removed more accurately than the water sampling method.
  • the conventional phantom calibration method has a problem that the number of sample points decreases as the pixel position of the detector approaches the periphery.
  • An object of the present invention is to provide a radiation tomography technique in a radiation tomography apparatus capable of reducing artifacts occurring in a reconstructed image with high accuracy and improving image quality of the reconstructed image.
  • the radiation tomography apparatus of the present invention has the following features.
  • a typical configuration example of the present invention will be described.
  • an imaging system including: a generation unit configured to generate radiation for irradiating a subject; a detection unit disposed to face the generation unit and configured to detect the radiation transmitted through the subject; A rotation unit for generating a tomographic image of the subject based on the plurality of transmission images captured at a plurality of rotation angle positions while rotating the imaging system around the subject.
  • a tomographic apparatus wherein each of a plurality of phantoms including at least one phantom whose cross sections perpendicular to the rotation axis of the imaging system have different sizes in two directions orthogonal to the rotation axis.
  • First storage means for storing three or more transmission images measured while rotating the imaging system around, and an image corresponding to the transmission image as a calculation image by calculation Generating means for generating, a second storage means for storing the generated calculated image, Correction means for correcting the intensity of the transmitted image of the subject based on the measured image and the calculated image.
  • the at least one phantom has a substantially elliptical cross section perpendicular to the rotation axis.
  • At least one of the plurality of phantoms has a substantially circular cross section perpendicular to the rotation axis, and The center is disposed at a position different from the rotation axis.
  • an imaging system including: a generation unit configured to generate radiation for irradiating the subject; a detection unit disposed to face the generation unit and configured to detect the radiation transmitted through the subject; A rotation unit for generating a tomographic image of the subject based on the plurality of transmission images captured at a plurality of rotation angle positions while rotating the imaging system around the subject.
  • a tomographic apparatus comprising: a plurality of phantoms including at least one phantom in which a cross section perpendicular to a rotation axis of the imaging system is substantially circular, and a center of the circle is disposed at a position different from the rotation axis.
  • a first storage means for storing three or more measured images of transmission images taken while rotating the imaging system around the image, and calculating an image corresponding to the transmission image by calculation Generating means for generating an image, second storage means for storing the generated calculated image, and correction for correcting the intensity of the transmitted image of the subject based on the measured image and the calculated image Means.
  • At least three of the plurality of phantoms have a substantially elliptical cross section perpendicular to the rotation axis.
  • at least one of the plurality of phantoms has a substantially circular cross section perpendicular to the rotation axis, and the center of the circle is the rotation center. It is characterized by being arranged at substantially the same position as the axis.
  • a center position of a tomographic plane of the phantom and the Phantom position calculating means for calculating a tilt in a parallel direction wherein the generating means determines a projection direction of the radiation at the time of creating the calculation image based on the center position and the tilt. It is characterized by doing.
  • the phantom position calculating means calculates a center position of a tomographic plane of the phantom based on a center of gravity of a signal intensity distribution of the tomographic image of the phantom. It is characterized by doing.
  • the phantom position calculating means is configured to calculate a position of the phantom in a direction parallel to a tomographic plane of the phantom based on a first-order approximation of a signal intensity distribution of a tomographic image of the phantom. It is characterized by calculating the inclination of.
  • An imaging system including: a generation unit configured to generate radiation for irradiating a subject; a detection unit disposed opposite to the generation unit to detect the radiation transmitted through the subject; and Rotating means for rotating around the object, generating a tomographic image of the object based on the plurality of transmission images taken at a plurality of rotational angle positions while rotating the imaging system around the object.
  • Radiation tomography apparatus comprising: a plurality of phantoms including at least one phantom having a cross section perpendicular to a rotation axis of the imaging system having different sizes in two directions orthogonal to the rotation axis.
  • First storage means for storing three or more transmission images measured around the phantom while rotating the imaging system, and generating an image corresponding to the transmission image as a calculation image by calculation Generating means for generating the calculated image; second storage means for storing the generated calculated image; and approximating the relationship between the signal intensity of the actually measured image and the signal intensity of the calculated image by an approximation function.
  • a parameter value deriving unit that derives a parameter value of a function
  • a third storage unit that stores the parameter value, and corrects the intensity of the transmitted image of the subject based on the measured image and the calculated image. Correction means.
  • An imaging system including: a generation unit configured to generate radiation for irradiating a subject; a detection unit disposed opposite to the generation unit to detect the radiation transmitted through the subject; and Rotating means for rotating around the object, generating a tomographic image of the object based on the plurality of transmission images taken at a plurality of rotational angle positions while rotating the imaging system around the object.
  • a radiation tomography apparatus having at least one phantom in which a cross section perpendicular to a rotation axis of the imaging system is substantially circular, and a center of the circle is arranged at a position different from the rotation axis.
  • first storage means for storing three or more transmission images measured around the phantom while rotating the imaging system, and calculating an image corresponding to the transmission image by calculation.
  • Generating means for generating a calculated image
  • second storage means for storing the generated calculated image, and approximating a relationship between a signal intensity of the actually measured image and a signal intensity of the calculated image by an approximation function.
  • Parameter value deriving means for deriving a parameter value of an approximate function
  • third storage means for storing the parameter value. Correcting the intensity of the transmitted image of the subject based on the measured image and the calculated image. And correction means for performing the correction.
  • the number of samples of measured data, which was at most 4 or 5 points in the conventional phantom calibration method, can be increased to 6 or more to several thousands, so that the accuracy of the phantom calibration method is improved.
  • the image quality of the reconstructed image can be improved.
  • FIG. 1 is a schematic front view of a radiation tomography apparatus according to Embodiment 1 of the present invention
  • FIG. 2 is a diagram for explaining the preprocessing means of the radiation tomography apparatus according to Embodiment 1 of the present invention
  • 3 is a diagram for explaining the structure of a data array in the preprocessing means of the radiation tomography apparatus according to the first embodiment of the present invention.
  • FIG. 4 is a correction table of the radiation tomography apparatus according to the first embodiment of the present invention.
  • FIG. 5 is a diagram for explaining a creating means
  • FIG. 5 is a diagram for explaining a correction table creating method of the radiation tomography apparatus according to the first embodiment of the present invention
  • FIG. 6 is a diagram for illustrating a first embodiment of the present invention.
  • FIG. 7 is a diagram for explaining signal intensity correction means of the radiation tomography apparatus.
  • FIG. 7 is a diagram for explaining an arithmetic method in a simulation image generation means of the radiation tomography apparatus according to the first embodiment of the present invention.
  • FIG. 8 shows a radiation tomographic image according to the first embodiment of the present invention.
  • FIG. 9 is a view for explaining a calibration phantom position detecting means of the imaging apparatus.
  • FIG. 9 is a view for explaining a method of arranging a plurality of elliptical phantoms of the radiation tomography apparatus according to the first embodiment of the present invention.
  • FIG. 0 is a diagram for explaining a calculation method in a simulation image generating means of the radiation tomography apparatus according to Embodiment 2 of the present invention
  • FIG. 11 is a radiation tomography apparatus according to Embodiment 2 of the present invention
  • FIG. 12 is a view for explaining an arrangement method of a plurality of cylindrical phantoms
  • FIG. 12 is a view for explaining another example of a calibration phantom of the radiation tomography apparatus according to the embodiment of the present invention.
  • FIG. 3 is a diagram for explaining an example of an image quality improvement effect by the radiation tomography apparatus according to the first embodiment of the present invention.
  • FIG. 1 is a schematic front view of a radiation tomography apparatus according to Embodiment 1 of the present invention.
  • the radiation tomography apparatus according to the first embodiment includes an X-ray tube 1, an X-ray detector 2, a rotating plate 4, a driving motor 5, a driving belt 6, a gantry 7, an imaging control unit 100, and a preprocessing unit 1.
  • 1 Measurement image memory 101, Correction table creation means 102, Simulation image memory 103, Signal strength correction means 104, Correction table memory 105, Simulation image generation means 106, image reconstruction means 107, calibration phantom position detection means 108, image display means 109, console 110, etc.
  • the X-ray tube 1 and the X-ray detector 2 are collectively referred to as an imaging system.
  • the imaging system is fixed to the rotating plate 4.
  • the drive motor 5 rotates the rotary plate 4 and the entire photographing system via a drive belt 6.
  • the imaging system irradiates the subject 3 with X-rays from all directions and captures the X-ray transmission image.
  • the rotation axis (not shown) of the rotating plate 4 is referred to as the Z axis.
  • the horizontal and vertical coordinate axes with the rotation center O of the rotating plate 4 as the origin are defined as the X axis and the Y axis, respectively.
  • the XYZ coordinate system defined by the X axis, Y axis, and Z axis is a rectangular coordinate system.
  • a typical example of the distance between the X-ray generation point S of the X-ray tube 1 and the rotation center O is 690 mm.
  • a typical example of the distance between the rotation center O and the X-ray input surface of the X-ray detector 2 is 380 mm.
  • a typical example of the time required for one rotation of the rotating plate 4 is 0.5 second.
  • the X-ray detector 2 As the X-ray detector 2, a known multi-slice X-ray detector including a ceramic scintillator and a photodiode is used.
  • the X-ray detector 2 is composed of a large number of detection elements (not shown).
  • a typical example of the number of the elements is 8996 in the XY plane direction (hereinafter referred to as a channel direction), and the Z-axis direction ( In the following, the slice direction is used).
  • Each detection element is arranged on an arc approximately equidistant from the X-ray generation point s, and a typical example of the input surface size is 1 mm in both the channel direction and the slice direction.
  • the radiation tomography apparatus has two types of imaging modes, a main imaging mode and a calibration imaging mode.
  • the examiner instructs the selection of the main imaging mode and the calibration imaging mode through the console 110.
  • broken arrows indicate the flow of data processing in the main shooting mode.
  • the solid arrows indicate the flow of data processing in the calibration imaging mode.
  • the imaging control means 100 starts the rotation of the rotating plate 4 via the drive motor 5.
  • the imaging control means 100 instructs the X-ray irradiation timing of the X-ray tube 1 and the imaging timing of the X-ray detector 2 and the entire circumference of the subject 3 Acquire shooting data from the direction.
  • the preprocessing unit 111 performs preprocessing including offset processing, air calibration processing, and logarithmic conversion processing on the photographing data by using a method described later, and the photographing data after the preprocessing is performed.
  • the signal strength correcting means 104 reads the measured image recorded in the measured image memory 101 and corrects the signal strength of the measured image using a method described later.
  • the signal strength correction means 104 refers to a correction table recorded in advance in the correction table memory 105, and performs a predetermined correction based on the correction table. The details of the correction table will be described later.
  • the image reconstructing means 107 reconstructs a radiation tomographic image of the subject 3 based on the output value of the signal intensity correcting means 104 by using a known technique. Finally, the reconstructed radiation tomographic image is displayed by the image display means 109.
  • the operation of the radiation tomography apparatus in the calibration imaging mode will be described.
  • a calibration fan to be described later is arranged as the subject 3.
  • photographing of the calibration phantom and pre-processing of the photographed data are performed in the same procedure as in the actual photographing mode. 0 Recorded in 1.
  • the actual measurement image for calibration is read out by the signal intensity correcting means 104.
  • the signal strength correction means 104 determines whether or not the correction table has already been recorded in the memory for correction table 105, and if it exists, refers to the correction table and refers to the correction table. Performs signal intensity correction on the measured image. If the correction table does not exist, the signal intensity correction to the calibration actual measurement image is omitted.
  • the image reconstruction means 107 reconstructs a radiation tomographic image of the calibration phantom based on the output value of the signal intensity correction means 104 by using a known technique.
  • the calibration phantom position detecting means 108 calculates the position of the calibration phantom on the XY plane using a method described later, based on the radiation tomographic image.
  • the simulation image generating means 106 uses a method described later to calculate the theoretical value of the photographing data of the calibration phantom (hereinafter referred to as a simulation image or a simulation image). Is calculated image), and the calculation result is recorded in the simulation image memory 103.
  • the correction table creating means 102 stores the calibration measurement images and the stains recorded in the measurement image memory 101. Based on the simulation image recorded in the simulation image memory 103, the method described below is used to create the correction table data for converting the signal intensity of the actually measured image into the theoretical value and correct the created result. Record in table memory 105. When the old correction table data is already recorded in the correction table memory 105, the old correction table data is overwritten with the new correction table data.
  • FIG. 2 is a diagram for explaining preprocessing means 111 of the radiation tomography apparatus according to the first embodiment of the present invention.
  • FIG. 3 is a diagram for explaining the structure of a data array in the preprocessing unit 111 of the radiation tomography apparatus according to the first embodiment of the present invention.
  • the processing procedure in the pre-processing unit 111 will be described with reference to FIGS. 2 and 3.
  • the X-ray detector 2 is a multi-slice detector, and has 896 and 32 detection elements in the channel direction and the slice direction, respectively, as described above.
  • the numbers of detection elements in the channel direction and the slice direction are generally represented as N and M, respectively.
  • 900 times of imaging are performed during one rotation of the imaging system.
  • the number of times of photographing is generally represented as K.
  • the detected signal is denoted by I nm (k).
  • the pre-processing means 11 is provided with three types of processing: offset image generation processing, air image generation processing, and air calibration processing. Of these, the offset image creation processing and the air image creation processing are selected at the time of offset image shooting and air image shooting performed prior to shooting of the subject 3. The air calibration process is selected when the subject 3 is photographed. The following three types of processing procedures Will be described in order.
  • the offset image creation process is a process of creating an arithmetic average of these K offset images obtained by offset image capturing (taken without X-ray irradiation by the X-ray tube 1). .
  • the offset image is sequentially overwritten in the frame memory 200 each time a photograph is taken.
  • the frame memory 200 has a data structure as shown in FIG. 3A, and stores N XM pieces of imaging data corresponding to one frame of the X-ray detector 2.
  • the averaging means 201 sequentially reads out the offset image and performs an averaging operation represented by (Equation 1).
  • Numeral 04 has a data structure as shown in FIG. 3 (B), and stores N XM average offset image data corresponding to one frame of the X-ray detector 2.
  • these averaging images are created for the K air images obtained by air image shooting (taking X-ray irradiation with the X-ray tube 1 without placing the subject 3). This is the processing to be performed.
  • the above air image is sequentially overwritten in the frame memory 200 each time a photographing is performed.
  • the averaging means 201 sequentially reads out the air image and performs an averaging operation represented by (Equation 2).
  • the average air image after offset correction obtained in 02 is recorded in the air image memory 205.
  • the air image memory 205 has a data structure as shown in FIG. 3C, and stores N X M average air image data corresponding to one frame of the X-ray detector 2.
  • the air calibration process performs air calibration on the captured image of the subject 3 to correct the spatial distribution of X-ray energy radiated from the X-ray tube 1 and uneven detection sensitivity of the X-ray detector 2. Processing.
  • the photographed image is sequentially overwritten in the frame memory 200 every time photographing is performed.
  • Equation 4 Note that the offset correction means 202 directly reads out the photographing data from the frame memory 200 without passing through the averaging means 201. In the above calculation, the average offset image recorded in the offset image memory 204 is referred to. At the same time when the offset correction of the captured image by the offset correction means 202 is completed, the air calibration means 203 reads the captured image after the offset correction, and 5) Perform the air calibration process represented by.
  • Equation 5 the average air image recorded in the air image memory 205 is referred to.
  • the photographed image after the air calibration obtained by the air calibration means 203 is recorded in the measured image memory 101.
  • the above series of processes in the air calibration process is repeatedly performed K times each time a captured image is recorded in the frame memory 200.
  • the measured image memory 101 has a data structure as shown in FIG. 3D, and stores N XMXK captured image data corresponding to K frames of the X-ray detector 2.
  • FIG. 4 is a diagram for explaining the correction table creating unit 102 of the radiation tomography apparatus according to the first embodiment of the present invention.
  • FIG. 5 is a diagram for explaining a method of creating a correction table of the radiation tomography apparatus according to the first embodiment of the present invention.
  • the processing procedure in the correction table creating means 102 will be described with reference to FIGS.
  • the correction table creating means 102 is used in the calibration shooting mode.
  • a calibration phantom described later is arranged as the subject 3, and imaging data of the calibration phantom is acquired.
  • the photographing data is recorded in the measured image memory 101 after the air calibration process is performed by the preprocessing means 111.
  • the data recorded in the measured image memory 101 and the simulated image memory 103 have the same data structure.
  • the method of calculating the simulation image J ' nra (k) will be described later.
  • the data of the measured image and the simulation image should match ideally, but do not actually match. This is because the input / output characteristics of the X-ray detector 2 have nonlinearity, and the energy of the X-rays radiated from the X-ray tube 1 has a spectral distribution. X-ray quality changes (beam hardening effect).
  • the correction table creating means 102 creates a correction table for correcting the non-linearity.
  • the relationship between the measured image J nm (k) and the simulated image J ' nra (k) can be approximated by a function such as a polynomial.
  • FIG. 5 shows an example of the above polynomial approximation.
  • the value of the measured image J nm (k) is first set on the horizontal axis, and the value of the simulation image J ' nm (k) is set on the vertical axis.
  • Equation 6 It should be noted that a known technique such as the least squares method is used for the above approximation.
  • the maximum order L of the polynomial function a predetermined value set in advance is used.
  • L needs to be 2 or more, but L is desirably 3 or more in order to accurately approximate the nonlinear component.
  • L the above polynomial has three coefficients a nm (l), an n (2), and a nm (3) from Equation 6.
  • the number of shots K must be 3 or more.
  • the number of times of photographing K needs to be L or more.
  • the least-squares approximation means 401 is realized by software processing using a dedicated arithmetic unit or a general-purpose arithmetic unit.
  • FIG. 6 is a diagram for explaining the signal intensity correcting means 104 of the radiation tomography apparatus according to the first embodiment of the present invention.
  • the photographed image of the subject 3 acquired in the main photographing mode is recorded in the actually measured image memory 101 after the above-described air calibration processing is performed by the preprocessing unit 111.
  • the polynomial calculating means 60 1 reads from the calibration table memory 105 the coefficients a nm (L), anra (Ll) of the polynomial for the detected pixel (n, m), Read the value of a nm (l) and record it in the buffer memory 602.
  • the polynomial calculation means 6001 calculates the imaging data J nm (k) after the air calibration processing and the coefficients a nra a), a nn (L-1), and a nra (l) of the polynomial , respectively.
  • the polynomial operation means 6001 is realized by software processing using a dedicated operation unit or a general-purpose operation unit.
  • FIG. 7 is a diagram for explaining a calculation method in the simulation image generating means 106 of the radiation tomography apparatus according to the first embodiment of the present invention.
  • FIG. 7 illustrates a method of generating a simulation image when an elliptical phantom 700 is used as a calibration phantom.
  • the rectangular coordinate system XYZ shown in FIG. 7A is a stationary coordinate system fixed to the gantry 7.
  • the X-ray generation point S rotates on the XY plane, and its rotation center coincides with the origin O of the XY Z coordinate system.
  • the XY plane intersects X-ray detector 2 at intersection line 702.
  • the elliptical phantom 700 has an elliptical column shape, and is arranged such that the column direction substantially matches the Z axis.
  • the elliptical phantom 700 is made of a substantially uniform material and density.
  • a typical example of the material of the elliptical phantom 700 is polyethylene, but another material such as ataryl may be used instead.
  • the outer dimensions of the elliptical phantom 700 in the major axis direction, minor axis direction, and column direction are shown.
  • Representative examples of 2a, 2b, and H are 350 mm, 200 mm, and 300 mm, respectively.
  • the line of intersection 700 of the elliptical phantom 700 and the XY plane has an elliptical shape.
  • the pq coordinate system is a coordinate system fixed to the ellipse phantom 700, and its origin is defined as the center ⁇ of the ellipse represented by the intersection line 700.
  • the p-axis and the q-axis are the major and minor axis directions of the substantially elliptical shape represented by the intersection line 701.
  • the ellipse phantom 700 is arranged such that the center point O 'of the ellipse is located near the origin O of the XYZ coordinate system, and that the .p-axis substantially coincides with the X-axis.
  • the position (0, x, O'y) of the ellipse center point o on the XY plane is not completely (0, 0). Also, the angle ⁇ between the p-axis and the X-axis cannot be completely zero.
  • the values of the above parameters (0'x, O'y) and ⁇ that define the position of the elliptical phantom 700 are automatically detected by the calibration phantom position detecting means 108 by a method described later.
  • Equation 7 Equation 7 ) where w nra (k) is the X-ray beam emitted from the X-ray generation point S at the k-th frame and incident on the detection pixel position (n, m) of the X-ray detector 2.
  • 0 3 be the transit distance in the ellipse phantom 700.
  • ⁇ p is the X-ray absorption coefficient of the elliptical phantom 700.
  • 0 k represents the rotation angle of the X-ray generation point S with respect to the X axis at the time of capturing the k-th frame, and is expressed by (Equation 13).
  • the simulation image generating means 106 is realized by software processing using a dedicated arithmetic unit or a general-purpose arithmetic unit.
  • FIG. 8 is a diagram for explaining the calibration phantom position detecting means 108 of the radiation tomography apparatus according to the first embodiment of the present invention.
  • the elliptical phantom 700 is arranged such that its center position ⁇ ′ substantially coincides with the center ⁇ of the ⁇ plane. However, it is not necessary to match the two with high accuracy, and it is sufficient that the difference between the two is within, for example, several cm. By allowing such rough placement accuracy, the number of man-hours required for placement of the calibration phantom can be reduced.
  • the calibration phantom position detecting means 108 automatically detects the amount of displacement of the elliptical phantom 700 in the above arrangement.
  • the displacement is defined by the center position 0 'of the elliptical phantom 700 and the inclination angle ⁇ of the elliptical phantom 700 in the major axis direction (p-axis direction) with respect to the X axis.
  • the value of ⁇ is c CT value binarization means 8 0 0 referenced by stain Yu Configuration image generating means 1 0 6
  • image reconstruction is first the CT reconfiguration images of elliptical phantom 700 Read from configuration means 107.
  • the signal value of the CT reconstructed image is represented as R (i, j).
  • CT value binarization means 8 00 then refers to the prerecorded threshold R t in the threshold memory 8 0 3 compares the value of R t and R (i, j). At this time, if R t R (i, j), the signal value of R (i, j) is rewritten to 1. If R (i, j) ⁇ R t , the signal value of R (i, j) is rewritten to 0.
  • the value of the threshold R t is an intermediate value between the CT value of CT values and the outer region of the elliptical phantom inside area is set in advance.
  • the CT reconstructed image binarized by the CT value binarization means 800 is an ellipse. It takes signal values 1 and 0 in the inner and outer regions of the circular phantom, respectively.
  • the binarized CT reconstructed image is read out by the center-of-gravity calculation means 81 and the inclination calculation means 802, respectively.
  • the center of gravity calculating means 800 is a means for calculating the position of the center of gravity of the elliptical phantom 700. The position of the center of gravity coincides with the center position O 'of the elliptical phantom 700, and is calculated by (Equation 14). .
  • Equation 14 where X is the position of the pixel (i, j) in the XY coordinate system.
  • the inclination calculating means 800 is a means for calculating an inclination angle ⁇ of the major axis of the elliptical phantom 7 to the X axis.
  • the calibration phantom position detecting means 108 is realized by software processing using a dedicated arithmetic unit or a general-purpose arithmetic unit.
  • FIG. 13 is a diagram for explaining an example of an image quality improvement effect by the radiation tomography apparatus according to the first embodiment of the present invention.
  • FIG. 13 (A) shows a method of arranging the evaluation subject 1300.
  • Fig. 13 (B) and (C) and (C) show the profiles on the Y-axis of the reconstructed image obtained when there is no calibration process.
  • the evaluation subject 1300 is a cylindrical water phantom having a diameter of 350 mm.
  • the tube voltage of the X-ray tube 1 was set to 120 kV, the tube current was set to 200 mA, and the other imaging conditions were the same as those described in Example 1.
  • the obtained profile 1301 had low uniformity, and the difference in CT value was a maximum of 59 HU.
  • the obtained profile 1322 had high uniformity, and the difference in CT value was 4.8 HU at the maximum. Therefore, it was confirmed that the present invention improved the derivation accuracy of the CT value of the reconstructed image, and improved the image quality.
  • the radiation tomography apparatus has been described.
  • the present invention is not limited to only the first embodiment, and can be variously modified without departing from the gist thereof.
  • the size of the elliptical phantom 700 is limited to one type, the force S, and a plurality of elliptical phantoms 700a to (!) Having different sizes as shown in FIG.
  • the elliptic phantoms 700a to 700d are assumed to have their center positions arranged, for example, near the rotation center O of the imaging system.
  • a correction table creation means 1 0 2 is a polynomial function approximation shown in (Equation 6) performed on the measured calibration image and the simulation image obtained for all the elliptic phantoms 700a to 700d, and the obtained coefficient a nn (L), a nra (L - 1), memory for ⁇ ⁇ ⁇ a nra (l) the ToTadashi table 1 0 5 It shall be recorded.
  • the radiation tomography apparatus according to the second embodiment of the present invention uses a cylindrical phantom 1000 as a calibration phantom instead of the elliptical phantom 700 used in the first embodiment.
  • the configuration of the radiation tomography apparatus other than the above is the same as that described in FIGS. 1 to 6 of the first embodiment, and a description thereof will not be repeated.
  • FIG. 10 is a diagram for explaining a calculation method in the simulation image generating means 106 of the radiation tomography apparatus according to Embodiment 2 of the present invention.
  • FIG. 1 ⁇ describes a simulation image when a cylindrical phantom 100 is used as a calibration phantom.
  • the rectangular coordinate system XYZ is a stationary coordinate system fixed to the gantry 7.
  • the X-ray generation point S rotates on the XY plane, and its rotation center coincides with the origin O of the XY Z coordinate system.
  • the XY plane intersects X-ray detector 2 at intersection line 1002.
  • the cylindrical phantom 1000 has a cylindrical shape, and is arranged such that the column direction substantially matches the Z axis.
  • the cylindrical phantom 100000 is made of a substantially uniform material and density.
  • a typical example of the material of the cylindrical phantom 1000 is polyethylene, but other materials such as ataryl may be used instead.
  • the outer dimensions in the circular and column directions of 1000 are represented by 2r and H, respectively.
  • Representative examples of 2r and H are 250 mm and 300 mm, respectively.
  • the intersection line 1001 between the cylindrical phantom 1000 and the XY plane has a substantially circular shape, and the center of the circle is 0.
  • the cylindrical phantom 1000 is located at a position (0'x, O'y) where the substantially circular center point o, is different from the rotation axis, that is, the origin o of the XYZ coordinate system.
  • the automatic detection of the center point O, by the calibration phantom position detecting means 108, can be realized by the same method as described with reference to FIG. However, when the cylindrical phantom 100 is targeted, it is not necessary to detect the inclination angle ⁇ as in the case of the elliptical phantom 700. Therefore, the calculation by the inclination calculating means 802 is omitted, and only the position of the center point o, detected by the center-of-gravity calculating means 801 is referred to by the simulation image generating means 106.
  • Equation 1 w nm (k) is the amount of X that is emitted from the X-ray generation point S and enters the detection pixel position (n, m) of the X-ray detector 2 in the k-th frame. This is the distance that the line beam 1003 passes through the cylindrical phantom 1000. Is the X-ray absorption coefficient of the cylindrical phantom 100.
  • the emission angle of the X-ray beam 103 in the channel direction is ⁇
  • the distance between the intersection line 1002 and the detection pixel position (n, m) is h nm
  • the X-ray generation point S and the rotation center O are If the distance between is represented by d and the distance between the X-ray generation point S and the input surface of the X-ray detector 2 is represented by D, the passing distance w nn (k) can be obtained by ( Equation 18).
  • the simulation image generating means 106 is realized by software processing using a dedicated arithmetic unit or a general-purpose arithmetic unit.
  • the present invention is not limited to only the second embodiment, and it goes without saying that various modifications can be made without departing from the scope of the invention.
  • the size of the cylindrical phantom 100 is limited to one type, but a plurality of cylindrical phantoms 1000 a to (!) Having different sizes as shown in FIG.
  • the cylindrical phantoms 100 a to d may have their outer peripheral positions set at For example, it is assumed that they are arranged so as to be substantially inscribed in the visual field 900 of the X-ray detector 2.
  • the correction table creating means 102 Is obtained by performing the polynomial function approximation shown in (Equation 6) on the measured calibration image and the simulation image obtained for all the cylindrical phantoms 1000 a to d, and obtains the obtained coefficient a nm (L), anra (L-1), ⁇ a nm (1) shall be recorded in the calibration table memory 105.
  • the elliptical phantom 700 and the cylindrical phantom 1000.0 were used as the calibration phantoms, respectively, but the types of the calibration phantoms are not limited thereto. Absent.
  • a calibration phantom we simulated the abdomen of the human body as shown in Fig. 12 (A); the abdominal phantom 1200 and the chest of the human body as shown in Fig. 12 (B).
  • a simulated chest phantom 1 201 may be used.
  • the chest phantom 1 201 has holes 1202 and 1203 for simulating the lung field in the human body.
  • the amount of scattered X-rays generated in the imaging of the calibration phantom approaches the scattered X-ray amount generated in the imaging of the subject 3, so the signal intensity Correction accuracy by the capturing means 104 can be improved.
  • Correction table using calibration measurement images and simulation images obtained for calibration phantoms (including phantoms whose cross-section perpendicular to the rotation axis is approximately circular and located at approximately the same position as the rotation axis) 105 may be created.
  • a radiation tomography technique for acquiring a large number of measurement data samples of a calibration phantom based on simple measurement in the phantom calibration method is realized.
  • the accuracy of the polynomial approximation in the phantom calibration method can be improved, and the image quality of the reconstructed image can be improved.

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Abstract

Un tomographe comprend un système d'imagerie comprenant un moyen générateur destiné à générer un rayonnement appliqué à un objet ainsi qu'un moyen de détection face au moyen de génération et adapté pour détecter le rayonnement transmis à travers l'objet. Le système d'imagerie est mis en rotation autour de fantômes comprenant au moins un fantôme dont la section transversale perpendiculaire à l'axe de rotation de l'axe d'imagerie a des dimensions différentes dans deux directions perpendiculaires à l'axe de rotation du système d'imagerie. Pendant la rotation, le système d'imagerie capture trois images de transmission de chaque fantôme ou davantage. Le tomographe est caractérisé en ce qu'il comprend également un premier moyen de stockage destiné à stocker les images effectivement mesurées des images de transmission, un moyen générateur destiné à générer des images correspondant aux images de transmission par calcul sous la forme d'images de calcul, un second moyen de stockage destiné à stocker les images de calcul générées et un moyen de correction destiné à corriger les densités des images de transmission de l'objet selon les images effectivement mesurées et les images de calcul.
PCT/JP2004/008379 2003-07-30 2004-06-09 Tomographe WO2005011502A1 (fr)

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JP2007185358A (ja) * 2006-01-13 2007-07-26 Hitachi Medical Corp Ｘ線ｃｔ装置
JP2007222599A (ja) * 2006-01-26 2007-09-06 Toshiba Corp Ｘ線ｃｔ装置、コンピュータプログラム及びファントム保持具
JP2011530372A (ja) * 2008-08-13 2011-12-22 コーニンクレッカ フィリップス エレクトロニクス エヌ ヴィ 理想的なアイソセントリックではない三次元回転型のｘ線スキャナシステムにおける、較正用ファントムに基づく回転中心探索アルゴリズムを用いたリング・アーチファクトの較正方法
JP2013546221A (ja) * 2010-09-30 2013-12-26 アナロジック コーポレイション 非線形データ取得
US9125655B2 (en) 2010-07-16 2015-09-08 California Institute Of Technology Correction and optimization of wave reflection in blood vessels
US9656009B2 (en) 2007-07-11 2017-05-23 California Institute Of Technology Cardiac assist system using helical arrangement of contractile bands and helically-twisting cardiac assist device
JP7467389B2 (ja) 2021-06-14 2024-04-15 富士フイルムヘルスケア株式会社 ファントムおよび放射線撮像装置、光子計数型検出器の較正方法

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JP6607256B2 (ja) * 2015-06-18 2019-11-20 株式会社島津製作所 放射線検出素子の感度補正方法および放射線断層撮影装置
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Publication number Priority date Publication date Assignee Title
JP2007125129A (ja) * 2005-11-02 2007-05-24 Hitachi Medical Corp Ｘ線ｃｔ装置
JP2007185358A (ja) * 2006-01-13 2007-07-26 Hitachi Medical Corp Ｘ線ｃｔ装置
JP2007222599A (ja) * 2006-01-26 2007-09-06 Toshiba Corp Ｘ線ｃｔ装置、コンピュータプログラム及びファントム保持具
US9656009B2 (en) 2007-07-11 2017-05-23 California Institute Of Technology Cardiac assist system using helical arrangement of contractile bands and helically-twisting cardiac assist device
JP2011530372A (ja) * 2008-08-13 2011-12-22 コーニンクレッカ フィリップス エレクトロニクス エヌ ヴィ 理想的なアイソセントリックではない三次元回転型のｘ線スキャナシステムにおける、較正用ファントムに基づく回転中心探索アルゴリズムを用いたリング・アーチファクトの較正方法
US9125655B2 (en) 2010-07-16 2015-09-08 California Institute Of Technology Correction and optimization of wave reflection in blood vessels
JP2013546221A (ja) * 2010-09-30 2013-12-26 アナロジック コーポレイション 非線形データ取得
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JP7467389B2 (ja) 2021-06-14 2024-04-15 富士フイルムヘルスケア株式会社 ファントムおよび放射線撮像装置、光子計数型検出器の較正方法

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